There are many useful drugs on the market today for which traditional means of administration are far from ideal. Bolus injections and oral unit doses typically result in a high initial systemic concentration of the active agent, in excess of the therapeutic concentration, which falls off over time and which will fall below the therapeutic concentration if another bolus is not timely administered. The result is that the ideal therapeutic concentration is not consistently maintained, there is a risk of toxicity associated with high systemic exposure to the drug, and the maintenance of a minimally effective concentration is dependent upon repeated administration at prescribed intervals. Patient compliance with a dosing regimen is difficult to ensure, especially where the course of therapy is long or of indeterminate or lifetime duration. There is a need for methods to deliver these drugs more effectively, so that therapeutic concentrations are maintained constantly in the tissues intended to be treated over an extended period of time, with minimal vulnerability to the vagaries of patient compliance, and ideally with minimal systemic exposure or exposure of uninvolved tissues and organs.
Modern drug discovery methods have led to the development of many drugs which are far more potent, yet have poorer solubility, than drugs developed through traditional medicinal chemistry methods. The development of these often-complex drugs has resulted in a need for methods to deliver such drugs more effectively and efficiently as well.
Extended-release and controlled-release drug delivery systems have been developed to address these needs. Implanted pumps and reservoirs, with various mechanisms for regulating release of drugs, were among the first solutions to be developed. A wide variety of polymeric matrices, permeated with drug substance, have also been developed which serve as implantable drug reservoirs. These polymeric implants gradually release drug over the course of days, weeks, or months as the contained drug diffuses through and out of the matrix and into the surrounding tissue. Three principal advantages provided by polymeric drug delivery compositions are:
(1) Localized delivery of drug. The product can be implanted directly at the site where drug action is needed and hence systemic exposure of the drug can be reduced. This becomes especially important for toxic drugs which are related to various systemic side effects (such as the chemotherapeutic drugs).
(2) Sustained delivery of drug. The drug is released over extended periods, eliminating the need for multiple injections or oral doses. This improves patient compliance, especially for drugs for chronic indications requiring frequent administration, such as replacement therapy for enzyme or hormone deficiencies, or for extended antibiotic treatments for such tenacious diseases as tuberculosis.
(3) Stabilization of the drug. The polymer matrix protects the drug from the physiological environment, particularly circulating enzymes, thereby improving stability in vivo. This makes the technology particularly attractive for the delivery of labile proteins and peptides.
For the reasons above, the use of drug-infused polymer implants as sustained-release drug delivery devices is now well established. One class of existing implants consists of preformed devices, ranging in size from matchstick-sized cylindrical rods such as the Norplant™ (levonorgestrel) and Zoladex™ (goserelin acetate) implants, to microspheres such as are sold under the trade name Lupron Depot™ (leuprolide acetate).
A major disadvantage of the macroscopic devices is their physical size. Implantation of Zoladex™ rods, for instance, requires the use of 14- or 16-gauge needles, and implantation of Norplant™ rods requires a surgical incision under local anesthesia, with similar subsequent procedures to replace and/or remove them. (The Zoladex™ rods are bioerodable, whereas Norplant™ implants are based on a non-bioerodable silicone.) Self-administration of such implants is not feasible, and the required intervention of trained medical personnel greatly raises the cost and inconvenience of such treatments.
Drug-containing polymer implants have been reduced in size by the expedient of grinding or milling a mixture of a drug substance and a gel-forming polymer at low temperature, as described in U.S. Pat. No. 5,385,738. The resulting powder is then suspended in a non-aqueous viscous solvent, such polyethylene glycol or a biocompatible oil, to obtain an injectable composition.
The size problem has similarly been overcome with microsphere implants, which can be administered (and self-administered where appropriate) by injection of an aqueous suspension of the microspheres. Lupron Depot™, for example, can be comfortably injected with a 22- or 23-gauge needle. Because microspheres are not retrievable from the body, they are necessarily based on bioerodable polymers. However, if an aqueous suspension of microspheres is stored for any length of time, the drug will diffuse from the particles into the aqueous phase, furthermore the bioerodable matrix itself is prone to hydrolysis in an aqueous environment. For these reasons, the injectable aqueous suspension must be prepared at the time of injection. A second disadvantage is the need for intramuscular injection. Finally, preparation of the microspheres is a complex process that is not easily carried out reproducibly and reliably, and regulatory validation of the manufacturing process can be a significant obstacle to commercialization of such products.
Another class of implants differing from pre-formed solid devices is injectable liquids. Upon injection, these are transformed in situ into solid implants. This class of implants is typified by compositions that transform from a drug-containing liquid phase to a drug-infused gel phase upon exposure to a physiological environment. Such in situ gelling compositions have several advantages: they can be readily and reliably manufactured by standard methods, they can be stored in the form of easily-injected liquids, they can be placed locally to achieve local delivery, and they can flow prior to gelling so as to fill voids and create a less-visible subcutaneous implant. In addition, a gelling, implant can serve as a scaffold for cellular colonization and tissue growth.
There are various changes in conditions that can trigger the gelling of an in situ gelling composition. Among these are changes in pH, osmolality, temperature, water concentration, and alterations in specific ion concentrations.
Temperature-sensitive in situ gelling compositions generally change from a sol to a gel when the temperature exceeds a critical solution temperature, which in the case of drug delivery systems must be reasonably close to body temperature. An example is the polyethylene oxide-polypropylene oxide block copolymer, sold under the trade name Pluronic™ F 127. A 25-40% aqueous solution of this material will gel at about body temperature, and drug release from such a gel occurs over a period of up to one week. Such compositions have the disadvantage that they must be carefully protected from premature gelling, through refrigerated storage, and no bioerodable polymer has yet been developed that undergoes a sol-gel transition at about body temperature.
A hydrogel whose drug release profile is both temperature- and pH-modulated has also been reported (T. G. Park, in Biomaterials 20:517-521 (1999)).
Another class of compositions form gels upon contact with water. For example, glycerol monooleate (GMO) containing a drug can be injected as a liquid lamellar phase, which upon injection and exposure to water forms a highly viscous cubic-phase hydrate. The drug is released from the cubic phase over the course of several days. An example of an injectable drug depot product based on GMO is the metronidazole dental gel formulation marketed under the trade name Elyzol™. Due to the high water content of the cubic phase, GMO formulations are prone to rapid drug release and are limited in duration of effect to no more than about five days.
There are very few biocompatible liquid crystal compositions that meet the requirement for phase transition to a sufficiently viscous state at physiological conditions. Polymers that precipitate upon contact with water, on the other hand, are numerous, and present a more versatile approach to the formulation of compositions that gel upon contact with water. Approaches based on in situ gelling compositions are described in U.S. Pat. Nos. 4,938,763, 5,077,049, 5,278,202, 5,324,519 and 5,780,044, all of which are incorporated herein by reference.
For example, the Atrigel™ drug delivery system consists of a bioerodable poly(DL-lactide-glycolide) (PLGA) copolymer (75:25 molar ratio) dissolved in N-methyl-2-pyrrolidone (NMP). Pharmaceuticals may be blended into this PLGA solution at the point of manufacture, or they may be added by the physician at the time of use. The liquid product is injected subcutaneously or intramuscularly through a small gauge needle, whereupon displacement of the NMP carrier with water in the tissue fluids causes the PLGA to precipitate, forming a solid film or implant. The drug incorporated within the implant is then released in a controlled manner as the polymer matrix erodes with time in the body. PLGA-based implants of this type can release drug over a period of several months. An example of a product employing this technology is the leuprolide acetate formulation marketed under the trade name Eligard™.
The Atrigel™ system uses N-methylpyrrolidinone (NMP) as a solvent for the PLGA copolymer. NMP is a water-miscible, low-molecular-weight and low-viscosity solvent that rapidly diffuses from the implant. Rapid solvent escape from the injected composition can lead to rapid and uneven precipitation of the polymer, shrinking of the implant, and local irritation or even necrosis due to exposure of tissues to a high local concentration of solvent.
The use of liquid polymers as solvents for in situ gel-forming compositions has been described in U.S. Pat. No. 5,607,686 and in U.S. application Ser. No. 10/169,012 (US 2003/0082234), corresponding to international patent application PCT/KR00/01508 (WO 01/45742). However, according to these patents, polyethylene glycol is not suitable as a solvent for PLGA.
In situ polymer-precipitation systems solve many of the problems associated with implants, but some difficulties remain. There is a need for in situ gelling drug delivery systems with improved properties, a simple preparation procedure and low toxicity of excipient.
PEGs have the advantage of solubilizing different drugs than NMP; in particular pegylated proteins can be expected to be more soluble and/or miscible in PEGs than in NMP. An additional advantage is that PEGs are available in different molecular weights and have different viscosities. In many instances it important to be able to control the viscosity of the injected gelling agent, which is not possible with NMP.